The present invention relates in general to circulatory assist devices, and, more specifically, to improved monitoring of the flow rate being output from an implanted pump unit.
Many types of circulatory assist devices are available for either short term or long term support for patients having cardiovascular disease. For example, a heart pump system known as a left ventricular assist device (LVAD) can provide long term patient support with an implantable pump associated with an externally-worn pump control unit and batteries. The LVAD improves circulation throughout the body by assisting the left side of the heart in pumping blood. One such system is the DuraHeart® LVAS system made by Terumo Heart, Inc., of Ann Arbor, Mich. The DuraHeart® system employs a centrifugal pump with a magnetically levitated impeller to pump blood from the left ventricle to the aorta. An electric motor magnetically coupled to the impeller is driven at a speed appropriate to obtain the desired blood flow through the pump.
A typical cardiac assist system includes a pumping unit, electrical motor (e.g., a brushless DC motor integrated in the pump housing), drive electronics, microprocessor control unit, and an energy source such as rechargeable batteries and/or an AC power conditioning circuit. The system is implanted during a surgical procedure in which a centrifugal pump is placed in the patient's chest. An inflow conduit is pierced into the left ventricle to supply blood to the pump. One end of an outflow conduit is mechanically fitted to the pump outlet and the other end is surgically attached to the patient's aorta by anastomosis. A percutaneous cable connects to the pump, exits the patient through an incision, and connects to the external control unit.
A control system for varying pump speed to achieve a target blood flow based on physiologic conditions is shown in U.S. Pat. No. 7,160,243, issued Jan. 9, 2007, which is incorporated herein by reference in its entirety. A target blood flow rate may be established based on the patient's heart rate so that the physiologic demand is met. The control unit may establish a speed setpoint for the pump motor to achieve the target flow. It is essential to automatically monitor pump performance to ensure that life support functions are maintained. One important performance measurement is the pump flow rate.
The actual blood flow being delivered to the patient by the assist device can be monitored either directly by sensors or indirectly by inferring flow based on motor current and speed. Because of the amount of space that would be required for a flow meter and the desire to maintain minimal size of an implanted device, the blood flow rate is usually estimated indirectly based on the current flowing through the motor and the rotational speed of the motor/impeller. Examples of systems using flow rate estimation include U.S. Pat. No. 7,033,147 to Yanai et al and U.S. Pat. No. 7,160,243 to Medvedev, both incorporated herein by reference.
A typical centrifugal pump employs a design which optimizes the shapes of the pumping chamber and the impeller rotating within the chamber so that the pump operates with a high efficiency. The pumping chamber has a curved volute shape around the impeller which increases in area as it nears the outlet. The flow into the outlet is substantially tangential to the outer edge of the impeller, and any reverse flow into the pump via the outlet would flow in opposition to the driven direction of the impeller.
The natural pumping action of a patient's heart is pulsatile. Since the assist device is working in conjunction with the beating of the patient's heart, it is subject to this pulsatile flow. As a result, there may be times during the heart cycle in which the assist pump experiences a backflow (i.e., reversal of the flow direction against the impeller rotation). The backflow has been a potential source of inaccuracy in flow rates determined by conventional flow estimation systems. For known pump designs, the motor current exhibits a minimum at a zero flow rate (i.e., as flow decreases toward zero the motor current decreases toward a minimum value which occurs at exactly a zero flow). If the flow decreases further (i.e., reverses to a backflow) then motor current begins to rise because the backflow is working against the impeller. Thus, a backflow condition may be indistinguishable from a small forward flow for estimates based on motor current when the flow is near a zero flow. It would be desirable to be able to unambiguously estimate at least small levels of backflow in order to better assess the pump performance and the physiological state of the patient.